An Electromyographic Study of the Hip Abductor Muscles as Subjects With a Hip Prosthesis Walked With Different Methods of Using a Cane and Carrying a Load.

Over the last 30 years, advances have been made in the design and methods of fixation of the total hip arthroplasty. As a result, total hip replacement has substantially improved the quality of life for people, most notably those with osteoarthritis.[1,2] Despite the success of total hip arthroplasty, aseptic premature loosening of the prosthetic hip remains the principal long-term complication following the implant, regardless of method of fixation.[3] Surgical revision following a failed hip arthroplasty has a less satisfactory long-term functional result than the initial operation.[4] Furthermore, a revision is more expensive, technically more complicated, and physiologically more stressful to the patient than a primary hip replacement.

Research suggests that premature loosening of the prosthesis is caused primarily by excessive mechanical wear and the formation of polyethylene debris that, in turn, stimulates osteolysis and a weakening of the bone.[6-8] Although no direct evidence exists, it may be logical to assume that excessive and repetitive forces from the hip abductor (HA) muscles may, in certain cases, be a contributing factor to excessive wear and subsequent premature loosening of a prosthetic hip. The reason for this hypothesis is that forces from the HA muscles during the stance phase of walking are responsible for a disproportionately large amount of the total force on the hip.[9-11] In efforts to reduce the wear on a prosthetic hip, physical therapists often recommend assistive devices, exercises, or methods of performing activities that limit the force demands on the HA muscles.[12-20] This form of advice, commonly referred to as "joint protection,"[21] may be particularly warranted when fixation of the implant is performed without cement or if the person has weakened bone or a history of complications that have required previous surgical revision.

In this report, I describe the third in a series of experiments designed to test selected forms of hip joint protection that purportedly reduce the muscular demands on the HA muscles following total hip arthroplasty. In all 3 studies, surface electromyography (EMG) was used as a noninvasive means of assessing the relative demands on the HA muscles overlying the prosthetic hip. In the 2 previous studies, I determined the relative EMG activity from the HA muscles as subjects walked while either carrying a load[13] or using a cane.[14] The purpose of this study was to determine the relative demands on the HA muscles when similar subjects carried a load while simultaneously using a cane in the other hand. This study may provide people who are at high risk of mechanical failure of their prosthetic hip with an additional strategy to reduce the demands on their HA muscles while walking.

Background

The main function of the HA muscles is to provide frontal-plane stability to the hip in the single-limb support phase of the gait cycle.[9, 12] This stability is achieved when the HA muscles produce a frontal-plane torque that equals the frontal-plane torque produced by body weight. The gluteus medius muscle is the largest of the abductor group, occupying about 60% of the total abductor cross-sectional area.[22] The gluteus minimus muscle is also an important HA muscle.[23] as are the tensor fasciae latae muscle, the piriformis muscle, and the anterior fibers of the gluteus maximus muscle.[24] Due to the difference in length of the moment arms available to the HA muscles and body weight,[25] the HA muscles must produce a force of about twice the body weight to ensure frontal-plane equilibrium. During the mid-stance phase of walking, this "myogenic" force adds to the joint compression caused by body weight, producing a total joint force equal to 3 to 3.5 times body weight.[26] Reducing the need for excessive force from the HA muscles, therefore, can be a logical way to minimize the forces produced across the prosthetic hip.

Previous research using EMG suggests that, when performed separately, carrying a load ipsilateral to a prosthetic hip or using a cane contralateral to a prosthetic hip each reduces the demands placed on the HA muscles.[13,14] As shown in the simplified static model shown in Figure 1A, a handheld load of 15% of body weight held ipsilateral to the prosthetic hip produces an external torque in the same rotary direction as that required by the right HA muscles (solid circles). As shown in Figure 1B, the upward-directed force from the cane held contralateral to the prosthetic hip also produces an external torque in the same rotary direction as that required by the right HA muscles (solid circles). In both cases, the external torques due to the load and the cane reduce the demand on the HA muscles. This reduced muscular demand was evident by a reduction in EMG activity from the HA muscles of the subjects with a prosthetic hip and an assumed reduction in force across the prosthetic implant.[13,14] When performed separately, therefore, carrying a load in the hand ipsilateral to the prosthetic hip and using a cane in the hand contralateral to the prosthetic hip are considered acceptable or preferred methods of hip joint protection. The degree to which combining these methods alters the demand on the HA muscles was the primary focus of this study. This is a logical question because carrying a load in one hand provides the opportunity to use a cane in the other hand.

[Figure 1 ILLUSTRATION OMITTED]

In the study described in this report, I focused on the biomechanical events during the mid-stance phase of walking, when, as depicted in Figure 1, the prosthetic hip is in single-limb support. The primary measurements were normalized surface EMG activity produced by the HA muscles and the force applied through the cane. The normalized EMG signal from the HA muscles was used as an index of the presumed relative force demands placed on the HA muscles. Based on the model shown in Figure 1, changes in HA muscle EMG activity were assumed to be associated with similar relative changes in abductor muscle-generated force across the prosthetic hip. Because the mathematical relationship between EMG activity and muscle force is not known for the HA muscles, I was not able to use EMG to directly reflect muscle or hip joint forces in this study.

The logic related to the model of Figure 1 and previous research[12-14,27] suggests that combining contralateral cane use with the ipsilateral load condition will offer reduction in EMG activity from the HA muscles. The primary hypothesis for this study, therefore, was that carrying a load ipsilateral to the prosthetic hip while using a cane contralateral to the prosthetic hip will generate less normalized HA muscle EMG activity than that produced while walking without carrying a load or using a cane. The secondary hypothesis was that carrying a load contralateral to the prosthetic hip while using a cane ipsilateral to the prosthetic hip will generate greater levels of normalized HA muscle EMG activity than that produced while walking without carrying a load or using a cane.

Method

Subject Selection Process

Twenty-four relatively active people with a unilateral hip prosthesis agreed to participate in this study. These subjects were also used in the prior 2 studies related to this topic,[13,14] and each subject was tested in one data collection session. Subjects were recruited through an advertisement placed in local newspapers, regional hospitals, and arthritis support groups. All subjects were paid for their time and received free consultation regarding their current exercise program and suggestions on how to minimize stress on their prosthetic hip.

Subjects selected for this study had to meet the following criteria. The prosthetic hip must have been in one hip only and not the result of a surgical revision. The operation must have been secondary to osteoarthritis; subjects with a hip implant due to rheumatoid disease, avascular necrosis, or congenital dysplasia were not selected. Furthermore, the prosthetic hip must have been the only joint prosthesis in the subject's body, and all other joints must have been relatively pain-free. Several subjects selected for the study reported they had minor joint pain, but they did not feel that their condition was disabling.

All subjects selected for this study declared they were in good health and independent in all activities of daily living. No subject reported respiratory disease, heart or vascular disease, or severe diabetes. All subjects reported not using an assistive device to aid in walking for distances less than 0.8 km (1/2 mile).

I conducted a brief physical examination of each subject and administered a brief questionnaire prior to each subject's acceptance into this study. The questionnaire was used to obtain personal data (eg, age, weight, duration of pain prior to surgery) and data related to the surgery (eg, side and date of surgery, use of cement). The examination consisted of checking for muscle weakness in the lower extremity (via manual resistance), hip instability, gait abnormality, pain, or other conditions that may compromise subject safety during the experiment. No subjects showed any of these conditions. Subjects selected for this study did not require any special footwear or orthosis for walking.

Subject Profile

Nine female subjects and 15 male subjects were used for this study. All subjects signed consent forms as required by Marquette University's Human Subjects Review Committee. Subjects ranged in age from 40 to 86 years ([bar] X=63.3, SD=10.7), in weight from 498.2 to 1,085.3 N(*) ([bar] X=757.5, SD=166.8), and in height from 1.52 to 1.91 m ([bar] X=1.71, SD=0.10), The time since hip surgery ranged from 5 to 96 months ([bar] X=24.9, SD=21.4). Twelve subjects had a prosthetic hip on their right side, and 12 subjects had a prosthetic hip on their left side.

Instrumentation

Surface EMG data were collected from the HA muscles as subjects walked on an indoor, hard-surfaced walkway. The EMG, footswitch, and cane instrumentation used in this study has been described in earlier work.[12,14,27] In brief, the EMG unit consisted of a surface on-site electrode, a ground electrode, an oscilloscope, a signal conditioning unit,([dagger]) a personal computer and analog-to-digital convertor, and software for data collection and data reduction. Raw bipolar EMG data were processed by using the root-mean-square (RMS) method by producing a linear envelope, or average voltage, over a specified time. The time constant used for the RMS processing was 55 milliseconds. The sampling rate of the processed EMG data was 100 times per second.

A calibrated electronic force transducer([double dagger]) was mounted in the shaft of a standard, aluminum adjustable cane,[sections] 17.8 cm (7 in) from the rubber tip.[14] The force transducer recorded the compression force produced parallel with the long axis of the cane. Subjects wore footswitches attached to galoshes that produced unique voltage levels to associate the EMG and cane voltage with a particular phase of gait. Two on-off footswitch closures defined the mid-stance phase of gait as the time interval between the instant of footflat and just prior to heel-off. The EMG, cane force, and footswitch voltages traveled between the subject and the signal processor and computer via a single 12.2-m (40-ft) cable.

Procedure

Pre-experimental protocol. Subjects were led to a private room for the application of the EMG electrodes, ground plate, and rubber galoshes. The skin over the posterolateral gluteal region was cleaned with alcohol. An EMG electrode was placed on the skin superficial to the belly of the gluteus medius muscle on the side of the prosthetic hip.[12,13] The ground electrode was placed over the anteromedial aspect of the tibia on the side of prosthetic hip. Proper electrode placement was verified by palpation of the gluteus medius muscle during isometric contraction and observing the raw EMG signal as the subject stood on one limb.

The height of the cane was adjusted to the appropriate height in the following manner. Subjects stood with a relaxed posture with the tip of the cane placed on the floor, 10.2 cm (4 in) lateral to the small toe. The height of the cane was then adjusted so that the elbow angle measured 30 degrees of flexion.[28] Subjects were shown how to carry a handheld load in the hand not occupied by the cane. The load was carried much like a person carries a single suitcase. The loads consisted of weights placed in a container with a hinged handle, with dimensions of 18 x 20 x 23 cm (7 x 8 x 9 in). Loads were adjusted to be 5%, 10%, and 15% of the subject's body weight.

Subjects were next asked to walk at a relatively constant self-selected walking speed. Pilot work indicated that the natural free walking speed of many subjects exceeded the speed that the heavier handheld loads could comfortably be carried. To reduce the self-selected walking speed, subjects were instructed to walk using a cane held in the hand opposite their prosthetic hip. (Using the cane had the desired effect of reducing walking speed to a level that coincided with a tolerable speed for carrying the heaviest loads.) After at least 3 minutes of walking, the subject's average walking speed was determined by stopwatch to the nearest 10th of a second ([bar] X=0.82 m/s [SD=0.09] for all 24 subjects). Each subject's average walking speed was determined over 3 trials of walking over a 10-m distance. Subjects repeated all subsequent walking trials at a speed within 10% of their own self-selected target speed. All subjects were able to learn to control their walking speed within the desired range.

Experimental protocol. Subjects practiced walking in a natural manner, with the various instrumentation in place, without using the cane or carrying a load. After confirmation of proper function of the instrumentation, the EMG amplifiers were set so the EMG voltage produced during a maximal isometric contraction of the HA muscles was well under the maximal limits expected by the computer's analog-to-digital convertor.

Before the start of the experiment, a pre-experimental EMG baseline, or control walk, was established for each subject. The HA muscle EMG baseline was determined by averaging the HA muscle EMG voltage produced during the mid-stance phase of walking at the subject's self-selected walking speed. During this normalization control walk, the subject walked at a self-selected walking speed without carrying a load or using a cane. For each control walking trial, the sampling of EMG data began as the subject walked across a 2-m mark on the walkway and continued for 10 seconds. Subjects were verbally instructed to stop walking after the 10-second period. An EMG baseline voltage was determined by averaging these data across the 4 control walking trials. The EMG voltages produced as subjects carried a load and used a cane were normalized to a percentage of this EMG baseline value (%EMG).

Each subject was exposed to 6 different experimental conditions, each involving a different combination of carrying 3 magnitudes of load while using a cane. The first set of conditions associated with the primary hypothesis had subjects use a cane held in the hand contralateral to their prosthetic hip while carrying a load weighing 5%, 10%, and 15% of their body weight positioned ipsilateral to their prosthetic hip. Subjects were instructed to carry the handheld load while placing the cane on the floor at the same time the limb with the prosthesis was in contact with the ground.

The second set of conditions associated with the secondary hypothesis had the same subjects walk while using a cane held in the hand ipsilateral to their prosthetic hip while carrying a load weighing 5%, 10%, and 15% of body weight contralateral to their prosthetic hip. As with the first set of conditions, subjects were instructed to carry the handheld load while pushing down on the cane at the same time the limb with the prosthesis was in contact with the ground. For all 6 conditions, subjects were instructed to push on the cane with a "moderate and comfortable level of force." Each subject participated in all 6 conditions performed at the subject's self-selected walking speed.

The order of the 6 conditions was randomized. Subjects were allowed a 90-second rest between walking trials as the experimenters displayed and verified the footswitch, EMG, and cane signal pattern on the computer screen. Data were accepted for analysis only after the target walking speed was confirmed and a typical footswitch pattern was displayed on the computer screen. Subjects performed a practice walk prior to each new condition.

Data were collected during the experimental trials in a similar method as that described for the pre-experimental baseline tests. One difference, however, was that only 2 walking trials of data were collected for each of the 6 conditions. This experimental design provided data on approximately 8 complete walking cycles per walking trial. On average, each of the 6 conditions produced data over 16 complete stance phases per subject.

Reliability Assessment of the HA Muscle EMG Data

Following the experiment, each subject reestablished a post-experimental baseline EMG value by repeating the pre-experimental control walking trials. This procedure allowed the experimenters to determine intrasubject reliability of the EMG baseline measurements by comparing the EMG voltages produced before and after the experiment. As a way to determine intrasubject reliability of the EMG measurements, a comparison was made between the grand mean value for EMG activity (in millivolts) produced in the pre-experimental control walking trials and the grand mean value for the EMG activity produced during the post-experimental control walking trials. Approximately 3 hours separated these 2 measurements. Each grand mean value was calculated by averaging all 24 subjects' EMG voltage from the side of the prosthetic hip during the mid-stance phase. The mean pre-experimental baseline EMG voltage was 147.4 mV, and the mean post-experimental baseline EMG voltage was 141.3 mV. This 4% difference in baseline EMG voltage was not considered to be significant. This difference would not have a systematic affect on the results of this study due to the random order of performance of the 6 cane and load conditions. A Pearson product-moment correlation coefficient of .993 and an intraclass correlation coefficient (ICC[1,k]) of .99 were calculated for the pre-experimental and post-experimental data (P [is less than] .0001).[29,30]

Data Analysis

The complete data set for all 24 subjects consisted of normalized EMG measurements (ie, %EMG) from HA muscles on the side of their prosthetic hip for each of the 6 cane and load conditions. All %EMG data were collected and averaged throughout the mid-stance phase of walking. Each mean %EMG value for each condition is based on a grand mean of approximately 16 complete gait cycles per subject, averaged over all 24 subjects.

A multifactorial analysis of variance (ANOVA) with a repeated-measures design was performed on the normalized %EMG data. The dependent variable was HA muscle %EMG produced on the side of the prosthetic hip, averaged over the mid-stance phase of walking. The independent variable was the cane or load condition. The HA muscle %EMG measurements for the 6 cane and load conditions were compared against each other and then against 0% (ie, the pre-experimental baseline EMG value) by using a multiple t test with Bonferroni adjustments.[13,44,29] These adjustments maintained the a priori alpha level by dividing .05 by the number of preplanned comparisons.

Results

The Table shows the descriptive statistics for the HA muscle %EMG and cane force for all 6 experimental conditions, averaged over all 24 subjects. An ANOVA performed on the mean %EMG produced by the HA muscles showed a main effect for the variable condition (F=141.8, P [is greater than] .0001). The %EMG mean produced for each of the 6 cane and load conditions is plotted in Figure 2. Each condition produced an HA muscle %EMG that was different from the 0% control baseline (ie, the EMG voltage produced during the control walk when subjects walked without a load or a cane). Furthermore, the mean HA muscle %EMG for the 3 contralateral cane and ipsilateral load conditions were equivalent to one another. In contrast, the mean HA muscle %EMG for the 3 ipsilateral cane and contralateral load conditions were different from one another.

Table. Descriptive Statistics for Hip Abductor (HA) Muscle %EMG(a) and Cane Force for All Six Cane-Load Conditions (N=24) During the Mid-stance Phase of Walking

                               [bar] X       SD
HA muscle %EMG
 CL cane and IL load(b)
   CL cane/IL load 5%          -36.6         19.3
   CL cane/IL load 10%         -39.8         19.1
   CL cane/IL load 15%         -46.1         16.0

 IL cane and CL load
   IL cane/CL load 5%           22.8         19.8
   IL cane/CL load 10%          50.6         25.4

   IL cane/CL load 15%          81.2         43.9

Cane force (N(c)) (force in
pounds shown in parentheses)

 CL cane and IL load(b)
   CL cane/IL load 5%          80.5 (18.1)   45.8 (10.3)
   CL cane/IL load 10%         76.5 (17.2)   43.1 (9.7)
   CL cane/IL load 15%         76.5 (17.2)   44.5 (10.0)

 IL cane and CL load
   IL cane/CL load 5%          64.9 (14.6)   47.1 (10.6)
   IL cane/CL load 10%         59.6 (13.4)   45.4 (10.2)
   IL cane/CL load 15%         52.9 (11.9)   43.1 (9.7)

                               Minimum       Maximum

HA muscle %EMG
 CL cane and IL load(b)
   CL cane/IL load 5%          -70.4           5.0
   CL cane/IL load 10%         -80.6           3.0
   CL cane/IL load 15%         -82.9         -16.0

 IL cane and CL load
   IL cane/CL load 5%          -30.0          57.0
   IL cane/CL load 10%           2.1         100.0
   IL cane/CL load 15%          14.0         179.0

Cane force (N(c)) (force in
pounds shown in parentheses)

 CL cane and IL load(b)
   CL cane/IL load 5%            8.9 (2.0)   187.7 (42.2)
   CL cane/IL load 10%          17.8 (4.0)   180.1 (40.5)
   CL cane/IL load 15%          16.9 (3.8)   177.5 (39.9)

 IL cane and CL load
   IL cane/CL load 5%           11.1 (2.5)   189.0 (42.5)
   IL cane/CL load 10%           6.2 (1.4)   163.2 (36.7)
   IL cane/CL load 15%          10.7 (2.4)    72.1 (16.2)

(a) %EMG = percentage of electromyographic voltage produced during the control walk (ie, while walking without a load or cane). Negative values indicate electromyographic activity less than that produced while walking a load or a cane. CL = contralateral, IL = ipsilateral.

(b) Load expressed as a percentage of the subject's body weight.

(c) 1 lb = 4.448 N.

[Figure 2 ILLUSTRATION OMITTED]

Discussion

Primary Hypothesis: Using a Cane in the Hand Contralateral to the Prosthetic Hip While Carrying a Load Ipsilateral to the Prosthetic Hip

As shown in the Table, all 3 combined contralateral cane and ipsilateral load conditions produced HA muscle %EMG means that were less than the baseline condition (Fig. 2). The primary hypothesis of this study, therefore, was accepted: carrying a load ipsilateral to the prosthetic hip while using a cane contralateral to the prosthetic generated less normalized HA muscle EMG activity than that produced while walking without a load or cane. The biomechanical model shown in Figure 3A explains, in part, the reason for the reduction in HA muscle %EMG compared with the control walk condition. The model assumes a person is in the single-limb support phase of gait, with the right prosthetic hip held in static equilibrium due to the balance of oppositely directed frontal-plane torques. (For simplicity, the frontal-plane torques produced by angular acceleration of the pelvis and trunk and load over the right prosthetic hip have been omitted from the model.) As shown in Figure 3A, the combined contralateral cane and ipsilateral load conditions produced torques that act in the same rotary direction as the torque produced by the right HA muscles. In this manner, the clockwise torque due to body weight (dashed circle) is offset by 3 counterclockwise torques (ie, that due to cane force, the weight of the load, and the force from the HA muscles [solid circles]).

[Figure 3 ILLUSTRATION OMITTED]

The marked reduction in HA muscle %EMG suggests that the 3-way force couple reduces the frontal-plane torque normally produced by the HA muscles. The reduced HA muscle %EMG is likely associated with a reduction in prosthetic hip reaction forces. Although the actual magnitude of this force cannot be determined from this study, a theoretical estimate can be made as shown in Figure 3B. The calculations are based, in part, on the average body weight of subjects in this study and the average cane force exerted while carrying a load weighing 15% of body weight. As shown, combining the "preferred" methods of using a cane and carrying a load requires only 195.9 N (44 lb) of HA muscle force, resulting in 876.9 N (197.1 lb) of prosthetic hip reaction force. In theory, this is only 46% of the prosthetic hip reaction force produced when a person with an identical anthropometric profile stands in single-limb support but not using a cane or carrying a load.[14] The reduction in prosthetic hip reaction is due primarily to a reduced demand on the HA muscles.

Figure 4 presents a summary of the %EMG data produced by the HA muscles from this study and 2 previous related EMG studies.[13,14] As depicted in the figure, the contralateral cane, the ipsilateral load, and the combined cane and load conditions generated equal or less %EMG from the HA muscles than the baseline EMG value. Although all conditions appear to offer hip joint protection, all 3 conditions probably are not equally beneficial when considering the force environment across both hips. Carrying a load ipsilateral to a prosthetic hip, either with or without using a cane, offers protection to the prosthetic hip at the "expense" of significant increase in HA muscular activity over the opposite hip.[13] Using the model shown in Figure 3A as a reference, the ipsilateral load held by the right hand becomes a contralateral load relative to the left hip as the left lower extremity cycles through the mid-stance phase. As previous studies have shown, a contralateral load increases the EMG activity of the HA muscles[12,13] and increases the pressure and force on the prosthesis.[19,31] Increasing the muscular load on the nonprosthetic hip may not be tolerated by people with marked bilateral hip disease. If the opposite (nonprosthetic) hip is healthy, however, then the combined contralateral cane and ipsilateral load condition can be recommended as an effective method of protecting the prosthetic hip. My recommendation, however, only applies for the person who cannot avoid carrying unilateral loads all together (which theoretically is the ideal situation) and when protection of the prosthesis or prosthesis-bone interface is especially warranted.

[Figure 4 ILLUSTRATION OMITTED]

The most effective and practical means of protecting the hip while walking appears to be to use only the cane held contralateral to the prosthetic hip and to avoid carrying external loads.[14,19] As shown in Figure 4, using the cane contralateral to the prosthetic hip (without a load) produced a 30% reduction in HA muscle EMG activity below the baseline value.[14] Although using the cane alone failed to match the reduction in %EMG provided by the combined cane and load condition (Fig. 4), using only the cane is the preferred method of joint protection when considering the average muscular demand across both right and left hips.

Secondary Hypothesis: Using a Cone in the Hand Ipsilateral to the Prosthetic Hip While Carrying a Load Contralateral to the Prosthetic Hip

Figure 2 shows that the combined ipsilateral cane and contralateral load conditions produced HA muscle %EMG means that were greater than for the control walk. The secondary hypothesis of this study, therefore, was accepted: carrying a load contralateral to the prosthetic hip while using a cane ipsilateral to the prosthetic hip generates greater levels of normalized HA muscle EMG activity than that produced while walking without a load or cane. These results were expected based on previous EMG-related studies[13,14] that tested the effect of the cane and load conditions on the HA muscles sepa-rately. No strategy that combines the ipsilateral cane and contralateral load appears to offer hip joint protection. Increasing the magnitude of the contralaterally held load (regardless of using a cane) progressively increases the demand on the HA muscles, likely producing very large forces over the prosthetic hip.[13] An in vivo peak pressure of nearly 7.96 mPa (1 mPa=145 1b/[in.sup.2]) was reported on the superior pole of a prosthetic hip as a subject walked while using a cane held ipsilateral and a 10% of body weight load held contralateral to the prosthetic hip.[19] The peak pressure was the highest reported across all experimental conditions of carrying a unilateral handheld load (10% of body weight), using a cane, or all combinations of these conditions. Repeated pressures of this magnitude may be tolerated safely in the intact healthy hip, but not necessarily in the prosthetic hip of an individual at high risk of loosening or fracture of the implant.

The relatively large %EMG means depicted on the right side of Figure 2 can be at least partially explained qualitatively by the model shown in Figure 5. The combined ipsilateral cane and contralateral load condition produces three concurrent clockwise torques (dashed circles) that must be matched by the one set of HA muscles. Because the HA muscles operate with a relatively small internal moment arm (D), the HA muscles must generate a very large force to ensure frontal-plane equilibrium. This force demand on the HA muscles has been estimated at nearly 3 times body weight when holding a contralateral load of 15% of body weight.[13] Hip abductor muscles that are generally atrophic or weakened by hip surgery or prolonged immobility may not be able to generate this level of force. As a consequence, people with weakened musculature may lack frontal-plane stability at the hip while in the mid-stance phase of gait.[21]

[Figure 5 ILLUSTRATION OMITTED]

Summary

Carrying loads (5%, 10%, and 15% of body weight) in the hand ipsilateral to a prosthetic hip while simultaneously using a cane in the hand contralateral to the prosthetic hip reduced the %EMG from the HA muscles overlying the prosthetic hip. This strategy offers at least the same level of reduction of HA muscle activity as performing each component separately. Combining these activities appears to be a method of protecting the joint only when carrying unilateral loads cannot be avoided. A more effective method of joint protection for the person at high risk for prosthetic hip failure is to avoid carrying unilateral loads altogether and to use a cane only in the hand contralateral to the prosthetic hip.

Carrying loads (5%, 10%, and 15% of body weight) in the hand contralateral to a prosthetic hip while simultaneously using a cane in the hand ipsilateral to the prosthetic hip increased the %EMG from the HA muscles overlying the prosthetic hip. This combined method of using a cane and carrying a load, in my opinion, should be avoided, especially by individuals who are at high risk of failure or loosening of their prosthetic hip. This study reinforces the finding of an earlier study[13] that carrying loads contralateral to the prosthetic hip produces all increase in %EMG from the HA muscles, and likely increases the force at the underlying prosthetic hip.[16,31]

Conclusion

The increase in average age of the population will likely result in a greater number of people receiving a hip replacement. Furthermore, data show that people who receive a prosthetic hip due to osteoarthrosis live longer than aged-matched controls.[32] The increased longevity is likely due to the fact that people who receive a prosthetic hip must be in relatively good health in order to tolerate the surgical procedure. Nevertheless, these factors suggest that, in the future, more people will be walking with a prosthetic hip for a greater number of years. In order to reduce a net increase in the number of required surgical revisions, continued advances are needed in the type of material used for the implant, surgical methods of fixation, and knowledge of how to limit excessive wear on the original prosthetic replacement.

Limitations of This Study

The model depicted in Figure 3 assumed a condition of static equilibrium over the prosthetic hip during the mid-stance phase of walking. For simplicity, the model excluded variables associated with the mid-stance phase, forces and torques outside of the frontal plane, and shifts in the center of mass while walking. In addition, the model assumed that all forces acted in the vertical direction. This model, therefore, contains error when estimating the absolute force and torque magnitudes. The model, however, provides a framework for understanding the approximate relative HA muscle-generated forces acting on the prosthetic hip during walking with a load or using a cane.

(*) 4.448 N=1 lb.

([dagger]) Therapeutics Unlimited, 2835 Friendship St, Iowa City, IA 52240.

([double dagger]) Genesco Technology Co, 650 Easy St, Simi Valley, CA 93065.

([sections]) Guardian: Sunrise Medical, 12899 Wentworth St, Arleta, CA 91331.

References

[1] Lieberman JR, Dorey F, Shekelle P, et al. Differences between patients' and physicians' evaluations of outcome after total hip arthroplasty. J Bone Joint Surg Am. 1996;78:835-838.

[2] Mancuso CA, Salvati EA, Johanson NA, et al. Patients' expectations and satisfaction with total hip arthroplasty. J Arthroplasty. 1997;12: 387-396.

[3] Kobayashi S, Takaoka K, Saito N, Hisa K. Factors affecting aseptic failure of fixation after primary Charnley total hip arthroplasty: multivariate survival analysis. J Bone Joint Surg Am. 1997;79:1618-1627.

[4] Robinson AH, Palmer CR, Villar RN. Is revision as good as primary hip replacement? A comparison of quality of life. J Bone Joint Surg Br. 1999;81:42-45.

[5] Barrack RL. Economics of revision total hip arthroplasty. Clin Orthop. 1995;319:209-214.

[6] Jasty M, Goetz DD, Bragdon CR, et al. Wear of polyethylene acetabular components in total hip arthroplasty: an analysis of one hundred and twenty-eight components retrieved at autopsy or revision operations. J Bone Joint Surg Am. 1997;79:349-358.

[7] Sieber HP, Rieker CB, Kottig P. Analysis of 118 second-generation metal-on-metal retrieved hip implants. J Bone Joint Surg Br. 1999;81: 46-50.

[8] Cuckler JM, Bearcroft J, Asgian CM. Femoral head technologies to reduce polyethylene wear in total hip arthroplasty. Clin Orthop. 1995; 317:57-63.

[9] Inman VT. Functional aspects of the abductor muscles of the hip. J Bone Joint Surg. 1947;29:607-619.

[10] McLeish RD, Charnley J. Abduction forces in the one-legged stance. J Biomech. 1970;3:191-209.

[11] Maquet P. Biomechanics of the Hip: As Applied to Osteoarthritis and Related Conditions. New York, NY: Springer-Verlag New York Inc; 1985.

[12] Neumann DA, Cook TM. Effect of load and carrying position on the electromyographic activity of the gluteus medius muscle during walking. Phys Ther. 1985;65:305-311.

[13] Neumann DA. Hip abductor muscle activity in persons with a hip prosthesis while walking carrying loads in one hand. Phys Ther. 1996; 76:1320-1330.

[14] Neumann DA. Hip abductor muscle activity as subjects with hip prostheses walk with different methods of using a cane. Phys Ther. 1998;78:490-501.

[15] Krebs DE, Elbaum L, Riley PO, et al. Exercise and gait effects on in vivo hip contact pressures. Phys Ther. 1991;71:301-309.

[16] Givens-Heiss DL, Krebs DE, Riley PO, et al. In vivo acetabular contact pressures during rehabilitation, part II: postacute phase. Phys Ther. 1992;72:700-705.

[17] Strickland EM, Fares M, Krebs DE, et al. In vivo acetabular contact pressures during rehabilitation, part I: acute phase. Phys Ther. 1992;72: 691-699.

[18] Fagerson TL, Krebs DE, Harris BA, Mann RW. Examining shibboleths of hip rehabilitation protocol using in vivo contact pressures from an instrumented hemiarthroplasty. Physiotherapy. 1995;81:533-540.

[19] McGibbon CA, Krebs DE, Mann RW. In vivo hip pressures during cane and load-carrying gait. Arthritis Care Res. 1997;10:300-307.

[20] Tackson SJ, Krebs DE, Harris BA. Acetabular pressures during hip arthritis exercises. Arthritis Care Res. 1997;10:308-319.

[21] Neumann DA. Biomechanical analysis of selected principles of hip joint protection. Arthritis Care Res. 1989;2:146-155.

[22] Clark JM, Haynor DR. Anatomy of the abductor muscles of the hip as studied by computed tomography. J Bone Joint Surg Am. 1987;69: 1021-1031.

[23] Kumagai MK, Shiba N, Higuchi F, et al. Functional evaluation of hip abductor muscles with use of magnetic resonance imaging. J Orthop Res. 1997;15:888-893.

[24] Dostal WF, Soderberg GL, Andrews JG. Actions of hip muscles. Phys Ther. 1986;66:351-359.

[25] Neumann DA, Soderberg GL, Cook TM. Comparison of maximal isometric hip abductor muscle torques between hip sides. Phys Ther. 1988;68:496-502.

[26] Rydell N. Forces acting on the femoral head-prosthesis: a study on strain gauge supplied prostheses in living persons. Acta Orthop Scand. 1966;37(suppl 88):1-132.

[27] Neumann DA, Cook TM, Sholty RL, Sobush DC. An electromyographic analysis of hip abductor muscle activity when subjects are carrying loads in one or both hands. Phys Ther. 1992;72:207-217.

[28] Kumar R, Roe MC, Scremin OU. Methods for estimating the proper length of a cane. Arch Phys Med Rehabil. 1995;76:1173-1175.

[29] SAS for the Personal Computer. 7th ed. Cary, NC: SAS Institute Inc; 1987.

[30] Shrout PE, Fleiss JL. Intraclass correlations: uses in assessing rater reliability. Psychol Bull. 1979;86:420-428.

[31] Bergmann G, Graichen F, Rohlmann A, Linke H. Hip joint forces during load carrying. Clin Orthop. 1997:335:190-201.

[32] Garellick G, Malchau H, Herberts P, et al. Life expectancy and cost utility after total hip replacement. Clin Orthop. 1998;346:141-151.

[33] Olson VL, Smidt GL, Johnston RC. The maximum torque generated by the eccentric, isometric, and concentric contractions of the abductor muscles. Phys Ther. 1972;52:149-158.

DA Neumann, PT, PhD, is Associate Professor, Physical Therapy Department, Marquette University, Box 1881, Milwaukee, WI, 53201-1881 (USA) ([e-mail]).

Dr Neumann, in addition to writing the article, provided concept and research design, data collection and analysis, project management, fund procurement, subjects and facilities/equipment (via Marquette University), and clerical support. The following people assisted with manuscript preparation, subject recruitment, data collection, graphics, or statistical support: John Rosecrance, PT, PhD, Richard Shields, PT, PhD, Thomas M Cook, PT, PhD, Richard Jensen, PT, PhD, Nick Schroeder, Tony Hornung, PT, Gregg Fuhrman, PT, and Mike O'Brien.

This study was approved by the Human Subjects Review Committee at Marquette University.

This research was presented, in part, at Physical Therapy '99: Annual Conference and Exposition of the American Physical Therapy Association; June 5-8, 1999; Washington, DC.

This project was funded by a grant from the National Arthritis Foundation.

This article was submitted May 5, 1999, and was accepted July 26, 1999.